The current state of the art of implantable cardiac assist technology consists of either a cannula-coupled constant flow pump which heavily supplements the left ventricle, or an implantable intra-aortic balloon.
Intra-Aortic Balloon Pumps (IABPs) such as those described in U.S. Pat. No. 6,210,318 and EP 0192574 (also published as U.S. Pat. No. 4,697,574) are well established technology which comprise of a mechanical device that sits at the top of the descending aorta. The device counterpulsates synchronously but in anti-phase with the heartbeat. This has the effect of increasing coronary blood flow and reducing afterload (known as aortic tension). The system works by mounting an inflatable polyethylene or silicone balloon on a catheter just distal to the left subclavian artery. A pneumatic (helium or carbon dioxide) line is then run through the arterial system, having entered the body through the skin (commonly at the groin femoral artery). During the hearts' diastolic (relaxing) phase the balloon is inflated, which serves to increase retrograde blood flow to the coronaries by displacing around 40 ml of blood from the balloon site, approximately half of which passess retrograde towards the aortic valve. Of this displaced blood, only 5 ml of displacement travels down the coronary arteries, with the rest perfusing the arteries that sit between the ascending and descending aorta. The retrograde coronary blood flow increases oxygen supply to the muscle of the heart (the myocardium). When the systolic (contracting) heart phase takes place the balloon deflates very quickly. This sudden release of pressure in the aorta offloads the ventricle, decreasing the work of the heart in systole and thereby the myocardial oxygen demand. This change in the myocardial oxygen supply and demand reverses the previous imbalance in the myocardium and thereby increases cardiac output.
Indications for an IABP include cardiogenic shock, intractable angina, and counteracting a low cardiac output after a coronary artery bypass graft (CABG). The IABP is amongst the most widely used bio-mechatronic human dynamic support systems.
Intra-aortic balloon pump insertion is traditionally performed through the femoral artery in the groin. However, this restricts the patient to bed rest, and prolonged implantation can be associated with infections in the groin crease and generalised sepsis from contamination of the balloon catheter from microorganisms present at the insertion site. Raman et al. (2010) describe a technique of insertion of a balloon pump through the subclavian artery, which allows the patient to ambulate. This technique can also be performed under local anaesthesia in the cardiac catheterization laboratory. IABPs cannot be deployed for long periods of time because of the risks of sepsis, arterial wall trauma, ischaemia in the relevant limb distal to the insertions site, thrombolic complications from exposure of the blood to a large surface area of foreign material, clotting abnormalities secondary to platelet consumption and device failure since their thin silicone balloons are prone to rupture.
In a chronic situation, where an intra-aortic balloon pump cannot be used (for the reasons described above), an extra-aortic balloon pump (EABP) may be used. EABPs such as those described in WO 02/24255, WO 2004/045677 and U.S. Pat. No. 4,733,652 address some of the problems associated with IABPs, being attached to the external surface of the aorta rather than being implanted within the aorta. These non-blood-contacting cuffs works in a similar manner to the IABP, though they surround the aorta and rely on external compression to displace blood. The insertion point through the skin can be variable (for example near the subclavian artery) due to the devices external arterial design. This allows for greater freedom of mobility for the patient, but the problem remains that a percutaneous tube is required (Legget et al, 2005). The EABP (Sunshine Heart C-Pulse) has not seen widespread clinical acceptance, where only a small number have ever been implanted worldwide. This is partly due to the restricted space available for installation around the ascending aorta. More significantly, the design depends on the natural ascending aorta to still be in place. At a time when such a device might be implanted, the aorta itself will tend to be diseased, becoming hardened through atherosclerosis and likely suffering from a narrowing of bore (stenosis). The compression of a hardened and constricted aorta has been found to be contraindicated.
Other disadvantages of EABPs can include increasing atheromous emboli (accumulation of plaque inside the aorta) through interaction with the aortic wall and migration and interference with neighbouring structures e.g. erosion of the pulmonary artery or lungs. That said, the placement around the ascending aorta means the device need only displace 10 ml of blood (5 ml in either direction) to achieve the equivalent effect as a 40 ml displacement of an IABP in the descending aorta. This is due to the retrograde losses into the three ascending arteries on the aortic arch being negated.
Other technologies which similarly augment blood flow using balloon pumps include a balloon pump for insertion into the descending aorta described in U.S. Pat. No. 6,030,335 but this device has a rigid outer body which precludes its implantation at the optimal position in the lower ascending aorta where pumping effect is optimised. Conduit mounted balloon pumps described in U.S. Pat. No. 4,195,623 and U.S. Pat. No. 4,015,590 are similarly non-optimally positioned. As far as the inventors are aware, these devices are not currently in clinical use.
In experimental studies by Furman et al., (1970, 1971) it was shown that diastolic counterpulsation is more effective at the level of the ascending aorta, for a number of reasons. Firstly, the closer the diastolic pulse wave generation is to the aortic valve, the more accurately counterpulsation can be timed to ventricular systole without conflict between the ventricular wavefront progressing distally from the aortic valve and the assist wave progressing proximally. Thus, proximity to the aortic valve allows better synchrony of the cardiac cycle to counterpulsation. Also, pulse propagation is minimized, reducing counterpulsation efficiency loss.
Substantial investigation has been carried out into different methods for cardiac assist devices. EABP and IABP have improved substantially in terms of reliability and thus increased patient survival rates. Aortic Balloon Pumps (ABP) are the most common mechanical circulatory assistance device used today (Dedhia et al, 2008). FIG. 1a illustrates the main components of the heart in a schematic cutaway view. FIG. 1b illustrates an IABP in situ within the aorta, showing the IABP inflated during the diastolic phase. FIG. 1c illustrated an EABP in situ around the aorta, showing the EABP inflated during the diastolic phase. By inflating and thus occluding (blocking fully or partially) the aorta during the diastolic (relaxing) phase of the heartbeat, the blood pressure is proximal to the balloon in the ascending aorta. This increases coronary blood flow. As the heart enters the systolic (contracting) phase, the balloon is released, which allows blood to rush into the aorta. This offloads the contraction of the heart as it displaces blood into the circulation. The pump therefore works by increasing coronary artery blood flow and reducing the afterload on the heart.
Other types of VAD include numerous types of pumps known as axial flow pumps, which work by increasing the blood follow in parallel with the natural heart. These tend to provide constant flow rather than being pulsatile in their pumping action. Problems with such devices include the pumps being solid and bulky, and not always easily fitting in the chest cavity of smaller human torso. Such pumps use a lot of power to rotate between 8,000-12,000 revolutions per minute, and so require bulky external battery packs. All feature similar power consumption (˜10 W), have percutaneous cables that protrude through the abdominal wall, and require ˜2.5 Kg control unit plus lithium-ion or lead acid batteries. The high speed impellors also have a tendency to damage red blood cells, or cause hemolysis (a separation of red blood cells and haemoglobin). A further problem is that regions of stagnant flow can result in thrombus formations.
There are numerous axial flow pumps (AFP), incorporating impellers of varying geometry. Amongst the most popular (Xinwei, 2003), at least in US markets, where figures are most readily available, are Medtronic Hemopump, Micromed DeBakey, Jarvik 2000 and Streamliner. These designs vary mostly around impeller and casing size, speed of rotation and impeller geometry. They generally work in parallel with the natural heart—the left ventricle still pumps in pulsatile fashion, while the AFPs output a continuous flow of 5-7 L/min at 13 kPa pressure. Variations in design include the Jarvik, which is small enough to sit inside the ventricle, making it quieter and lighter than others. The Streamliner features a magnetic levitation impeller which reduces friction, thus heating the blood and also reduces non-encapsulated (wetted) components. It also increases complexity of control significantly, requiring constant monitoring and shifting of the magnetic field, to keep the impeller aligned. The DeBakey has a specially designed low shear geometry (designed in conjunction with NASA).
There are a number of negative responses from the body arising from the use of axial flow pumps. These stem from the introduction of a constant blood flow. Kidney function can deteriorate, reflexes are reduced and the function of the endocrine system is sometimes altered. These devices are designed to run in parallel with the heartbeat, thus a small pulsatile flow is maintained. While this serves to reduce these issues, full pulsatile flow would be preferable.
With all types of ventricular assist devices, the machines used to support such intervention either cannot be implanted inside the body or can only be implanted for a few days. The therapy also inevitably involves the patient being bedbound in hospital whilst being attached to a machine that is partially implanted into the patient, involving the need for transdermal lines and catheters. On average there are 754.5 hospital bed days per patient utilised from primary diagnosis (Sutherland et al, 2010). With an ever increasing risk of hospital born infections such as MRSA and C. Diff, there is an absolute need to design a fully implantable system which will give the patient the same level of mechanical heart assistance without needing prolonged periods confined in a hospital environment. Such a system is not currently available.
It is an object of the invention to address one or more of the above mentioned problems.